The invention relates generally to nanostructured devices for detecting or analyzing biomolecules and their interactions.
Proteomics offers great potential for discovering biomarker patterns for earlier screening and detection of lethal and infectious diseases, systematic monitoring of physiological responses to drugs, and selecting the best treatment options for individual patients. For routine clinical use, an inexpensive, easy-to-use, multiplexed and high throughput protein analysis platform is needed, with high sensitivity and specificity for detection of low-abundance biomarkers in serum or other body fluids. There is also a need for high throughput and highly integrated sensor arrays for drug screening.
Nanostructured sensor arrays that use purely electrical detection, such as a field effect transistor (FET), fabricated with Si or other semiconductors, offer some of the desired characteristics. In such a device, a device channel of Si or other semiconductors is defined between two electrodes. The surface of the semiconductor channel or its oxide surface may be modified and covalently functionalized with antibodies or other receptor ligands for quantitative biorecognition. The binding of protein or other biomolecules induces net charge change, or change in dipole moment and binding-induced dipoles or modification of energy distribution and/or density of surface states. These binding events can change surface potential of the FET device and therefore modulate the conductance of the semiconductor channel. A small voltage or current, small enough not to disturb biomolecule interactions, is applied between two electrodes, and the change in conductance of the device channel is related and calibrated to the analyte concentration in a solution. When the device channel is reduced to nanoscale, the detection limit can be significantly reduced due to increased surface-to-volume ratio. Further, the response time can also be reduced due to favorable mass transport at low analyte concentrations due to small binding capacity of the small sensing surface. The ultralow detection limit of the nano-FET sensor at low ionic strength solutions has been recently demonstrated.
However, these devices may be rendered ineffective due to the screening effect in higher ionic strength solutions. The Debye screening length is defined as the distance from the sensing surface where potential change can be detected by the sensing device. In a high ionic strength solution, the screening length is reduced by ions and thus, analytes present beyond the screening length cannot be detected. As shown in FIG. 1, the Debye screening length decreases with an increase in ionic strength, therefore the binding events may not be detectable in high ionic strength solutions. It would be desirable to provide a method and a device that would enable a nano-FET biosensor to operate at higher ionic strengths when physiological samples with high ionic strength are to be analyzed, such as analyzing protein biomarkers in serum or other body fluids.
The Debye-Huckel Theory is useful to better understand the issues associated with operating biosensing devices in higher ionic solutions. For example, assuming a perfect orientation of an immobilized antibody, FIG. 2 shows the interaction between an antigen and an antibody in a solution. The potential distribution as a function of distance away from the electrolyte-immuno FET interface with immobilized antibodies is shown for both high ionic strength and low ionic strength cases. It can be seen from FIG. 2 that the electrostatic potential decreases rapidly as the binding site moves away from the electrolyte/gate insulator interface. The Debye screening length, δ, can be simply defined, in this example, as the distance away from the electrolyte/gate insulator interface at which a charge redistribution can still be detected by the FET sensor. In the high ionic strength environment, the Debye length is extremely short due to charge screening of the analyte antigen by excess ions (or more precisely “counterions”) present in solution. From the FET perspective, this charge screening effect makes it ineffective to detect charges induced by antigen/antibody interaction beyond the screening distance and therefore the antigen molecule must come closer to the sensor surface in order for its intrinsic or induced charges to be detected. Beginning with a buffer solution with an ionic strength solution of 0.2M, for example, the calculated Debye length is approximately 1 nm, which is significantly shorter than the average length of an antibody molecule (˜10 nm). Therefore the binding of antigen to the antibody receptor results in the redistribution of charges too distant to be detected by the FET sensor. However, in the absence of such excess charged species in solution, as in the case for low ionic strength situation, the screening effect by counterions is not as severe. The Debye length is much longer and the antigen molecule can be detected at a distance that is further away from the sensing surface. The overlapping of potentials in FIG. 2 in the low ionic strength case signifies a measurable effect with potentiometrically-operated immunoFET. The equation for Debye screening length in electrolytic solution is illustrated as:
  δ  =                    ɛ        ⁢                                  ⁢                  K          B                ⁢        T                    8        ⁢        π        ⁢                                  ⁢                  e          2                ⁢        I            where KB is the Boltzmann constant, T is temperature, e is the elementary charge (1.6×10−19 C), ∈ is the dielectric constant, and I is the ionic strength which has the expression
  I  =                    ∑        1        i            ⁢                        n          i                ⁢                  Z          i          2                      2  where ni represents the concentration of the ith ionic species in the electrolytic solution and Zi is the charge of the ith species. The sum of the product of the concentration and charge of all ionic species gives an estimate of the ionic strength of the electrolytic solution. Since the Debye length varies as the inverse square root of the ionic strength, the sensing response depends on the ionic strength of the solution.
The nanoscale channel can increase surface-to-volume ratio of the device and therefore significantly lower the detection limit, but lithography tools that are expensive and lower throughput are required to define nanoscale patterns. It would also be desirable to increase surface-to-volume ratio of the channel without reducing the channel to nanoscale size, so larger channel size can be used to achieve the low detection limit and more conventional and inexpensive lithography tools can be used. It can significantly reduce the cost of device fabrication.